Radiographic imaging such as x-ray imaging has been used for years in medical applications and for non-destructive testing.
Normally, an x-ray imaging system includes an x-ray source and an x-ray detector array consisting of multiple detectors comprising one or many detector elements (independent means of measuring x-ray intensity/fluence). The x-ray source emits x-rays, which pass through a subject or object to be imaged and are then registered by the detector array. Since some materials absorb a larger fraction of the x-rays than others, an image is formed of the subject or object.
An example of a commonly used x-ray imaging system is an x-ray computed tomography, CT, system, which may include an x-ray tube that produces a fan- or cone beam of x-rays and an opposing array of x-ray detectors measuring the fraction of x-rays that are transmitted through a patient or object. The x-ray tube and detector array are mounted in a gantry that rotates around the imaged object. An illustration of a fan beam CT geometry is shown in FIG. 3.
The dimensions and segmentation of the detector array affect the imaging capabilities of the CT system. A plurality of detector elements in the direction of the rotational axis of the gantry, i.e. the z-direction of FIG. 3 enables multi-slice image acquisition. A plurality of detector elements in the angular direction (ξ in FIG. 3) enables measurement of multiple projections in the same plane simultaneously and this is applied in fan/cone-beam CT. Most conventional detectors are so called flat-panel detectors, meaning that they have detector elements in the slice (z) and angular (ξ) directions.
X-ray detectors made from low-Z materials need to have a substantial thickness in the direction of the x-ray beam in order to have sufficient detection efficiency to be used in CT. This can be solved by, for example, using an “edge-on” geometry, as in U.S. Pat. No. 8,183,535, in which the detector array 50 is built up of a multitude of edge on detectors 5, which comprise thin wafers of a low-atomic number material, oriented with the edge towards the impinging x-rays 45. FIG. 2 shows a schematic illustration of an array of edge-on detectors 5, showing the position of the x-ray source 60, the direction of the x-rays 45, the detector array 50, a single edge-on detector 5 and the angular direction 55 of movement of the detector array 50. It is common that each photon-counting edge-on detector 5 has a plurality of detector elements 15 on a 2D grid on the wafer. An example of an edge-on semiconductor wafer is illustrated in FIG. 1, which shows the different detector elements 15 in a column on the detector array 50 and the direction of the impinging x-rays 45. Each individual wafer is, for example, oriented such that it has detector elements 15 in the slice direction (z) and in the direction of the x-rays 45, as schematically illustrated in FIG. 3. The edge-on geometry for semiconductor detectors is also suggested in U.S. Pat. Nos. 4,937,453, 5,434,417, US 2004/0251419 and WO 2010/093314. Wafer detectors that are oriented with a slight angle with respect to the direction of the x-rays 45 are normally also included in the term “edge-on”.
Detector elements 15 at different depths into the detector material with respect to the impinging x-rays 45 will be referred to as different “depth segments”. The detector elements 15 at different depths are generally aligned such that several detector elements 15 (from different depths) measure the same X-rays 45.
FIG. 9 is a schematic diagram illustrating a semiconductor detector module implemented as a multi chip module similar to an exemplary embodiment U.S. Pat. No. 8,183,535. In this example, the detector elements 15 are organized in three depth segments with respect to the direction of the incoming x-rays 45. This example shows how the semiconductor detector module also can have the function of substrate in a Multi Chip Module (MCM). The signal is routed 37 from the detector elements 15 to inputs of parallel processing circuits (e.g. ASICs) 30. It should be understood that the term Application Specific Integrated Circuit (ASIC) is to be interpreted broadly as any general circuit used and configured for a specific application. The ASIC processes the electric charge generated from each x-ray and converts it to digital data, which can be used to obtain measurement data such as a photon count and/or estimated energy. The ASICs are configured for connection to digital data processing circuitry 20 so the digital data may be sent to further digital data processing circuitry 20 and/or memories located outside of the MCM and finally the data will be the input for image processing to generate the reconstructed image.
For a given rotational position, each detector element 15 measures the transmitted x-rays 45 for a certain projection line. Such a measurement is called a projection measurement. The collection of projection measurements for many projection lines is called a sinogram. The sinogram data is utilized through image reconstruction to obtain an image of the interior of the imaged object. Each projection line (a point in the sinogram) is given by an angular coordinate, θ, and a radial coordinate, r, as defined in FIG. 7. Each measurement with a detector element 15 at a specific coordinate given by (r,θ) is a sample of the sinogram. More samples in the sinogram generally lead to a better representation of the real sinogram and therefore also a more accurately reconstructed image. An example of how a detector array 50, similar to that displayed in FIG. 3, samples the sinogram space is shown in FIG. 8 A for two angular positions of the gantry separated by Δθ. The different r positions of the samples come from the different detector elements 15 in the detector array 50.
Generally, the gantry rotates continuously and each detector element 15 measures the x-rays 45 within a frame time. A measurement period is here defined as the interval in time during which a certain detector element 15 is occupied with a measurement. The length of the measurement period can be, but does not have to be, equal to the frame time. The measurement period is much smaller than the total data acquisition time and multiple measurement periods follow directly after each other throughout the measurement. The length of the measurement period is referred to as the temporal sampling interval and the reciprocal of the sampling interval 1/T is referred to as the sampling frequency. The angular sampling interval of the CT system 10 is given by the angular velocity of the gantry, ω=dθ/dt, and the temporal sampling interval, T, by Δθ=ωT. A schematic illustration of the angular sampling is displayed in FIG. 4, where the photon counting edge-on detector 5 and the X-ray source 60 are illustrated for two different positions separated in time by the sampling interval T. The radial coordinate for all projection lines corresponding to a specific detector element 15 is invariant to the rotation of the gantry.
In order to perform an accurate image reconstruction from tomographic data, it is essential that there are a sufficient amount of angular samples. Insufficient angular sampling can lead to artifacts in the image, aliasing and poor resolution.
One way to increase the angular sampling frequency (without using specific oversampling schemes, as described in a later) is to decrease the temporal sampling interval T. This is, however, often limited by the detector electronics. Another way to obtain higher angular sampling frequency is to decrease the rotation speed ω and lower the flux (in order not to increase the patient dose). This comes with a noise penalty for conventional energy integrating detectors since less flux implies more relative electronic noise when integrating the signal. For photon counting detectors, however, decreasing the flux does not come with a noise penalty, since there is no integration process. Therefore, it is possible to use a higher sampling rate in photon-counting CT compared to conventional CT.
There are several oversampling schemes developed for computed tomography, for example: “quarter-detector offset” and “flying focal spot”. The “quarter-detector offset” method is well known and implies that the detector elements are spatially offset with respect to the central line of the fan-beam by one quarter of the detector width. This implies that the projections at θ and θ+180 degrees are not the same, but offset by half a pixel. This produces an oversampling (two times higher) in the radial direction. The method “flying focal spot” implies that the focal spot is moved during the measurement in order to produce more projection lines. This method can produce an oversampling in both the radial and the angular directions. For edge-on detectors, the flying focal-spot method has the disadvantage that the spectral response of the detector changes if you change the alignment of the detector with respect to the source by moving the source.
In U.S. Pat. No. 7,696,481 there is described a method for oversampling using for multi-layer detectors where the detector elements in the different layers are spatially offset with resect to each other. This produces an oversampling in both the radial and the angular direction. However, when low-Z materials are used as detector material, the fraction of photons which scatter in the detector is significant, therefore it can be beneficial to have anti-scatter modules interfolded between at least a subset of the detector modules, as described in U.S. Pat. No. 8,183,535 B2. If such anti-scatter modules are used, it is preferable to align the anti-scatter modules (and the detectors modules) with the direction of the x-rays 45 in order to maintain detection efficiency (if not aligned, the anti-scatter modules will also absorb primary radiation which otherwise could be collected by the detector). Therefore, if anti-scatter modules that are interfolded between the detector modules are used, an oversampling scheme that includes spatial shift between the detector elements in the different depth segments is impractical.